X-ray source with flexible probe

ABSTRACT

This invention is directed to an x-ray source comprising a power supply, a flexible fiber optic cable assembly, a light source, and a target assembly. The power supply includes a first terminal and a second terminal, and elements for establishing an output voltage between the first terminal and the second terminal. The flexible fiber optical cable assembly has an originating end and a terminating end, and includes a fiber optical element extending from the originating end to the terminating end. The cable is adapted for transmitting light incident on the originating end to the terminating end. The light source includes elements for generating a beam of light at and directed to the originating end of the fiber optical cable assembly. The target assembly is affixed to the terminating end of the fiber optical cable assembly and is electrically coupled to the power supply by way of the first terminal and the second terminal. The target assembly includes elements for emitting x-rays in a predetermined spectral range, in response to light transmitted to the terminating end.

REFERENCE TO RELATED APPLICATION

The subject matter of this application related to that of U.S. patentapplication Ser. No. 955,494, entitled LOW POWER X-RAY SOURCE WITHIMPLANTABLE PROBE FOR TREATMENT OF BRAIN TUMORS, filed Oct. 2, 1992; andU.S. Pat. No. 5,153,900, entitled MINIATURIZED LOW POWER X-RAY SOURCE,filed Sep. 5, 1990, the contents of both of which are incorporatedherein by reference. The subject matter of this application is alsorelated to that of U.S. patent application Ser. No. 08/184,271, entitledX-RAY SOURCE WITH SHAPED RADIATION PATTERN, filed on even date herewithand U.S. patent application Ser. No. 08/184,021, entitled X-RAY SOURCEWITH IMPROVED BEAM STEERING, filed on even date herewith.

BACKGROUND OF DISCLOSURE

The present invention relates to a miniaturized, low power, programmablex-ray source for use in delivering low-levels of substantially constantor intermittent x-rays to a specified region.

Conventional medical x-ray sources are large, fixed position machines.Generally, the head of the x-ray tube is placed in one room and thecontrol console in an adjoining area, with a protective wall, equippedwith a viewing window, separating the two. The x-ray tube typically isapproximately 20 to 35 centimeters (cm) long, and approximately 15 cm indiameter. A high voltage power supply is housed within a containerlocated in a corner of the room containing the x-ray tube. Patients arebrought to the machine for diagnostic, therapeutic, or palliativetreatment.

Diagnostic x-ray machines are typically operated at voltages below 150kilovolts (kV), and at currents from approximately 25 to 1200 milliamps(mA). By contrast, the currents in therapeutic units typically do notexceed 20 mA at voltages which may range above 150 kV. When an x-raymachine is operated at nominal voltages of 10 to 140 kV, the emittedx-rays provide limited penetration of tissue, and are thus useful intreating skin lesions. At higher voltages (approximately 250 kV), deepx-ray penetration is achieved, which is useful in the treatment of majorbody tumors. Super voltage machines, operable in the 4 to 8 megavolt(MV) region, are used to ablate or destroy all types of tumors, exceptsuperficial skin lesions.

A conventional x-ray tube includes an anode, grid, and cathode assembly.The cathode assembly generates an electron beam which is directed to atarget, by an electric field established by the anode and grid. Thetarget in turn emits x-ray radiation in response to the incidentelectron beam. The radiation absorbed by a patient generally is thatwhich is transmitted from the target in the x-ray tube through a windowin the tube, taking into account transmission losses. This windowtypically is a thin section of beryllium, or other suitable material. Ina typical x-ray machine, the cathode assembly consists of a thoriatedtungsten coil approximately 2 mm in diameter and 1 to 2 cm in lengthwhich, when resistively heated with a current of 4 amps (A) or higher,thermionically emits electrons. This coil is surrounded by a metalfocusing cup which concentrates the beam of electrons to a small spot onan opposing anode which also functions as the target. In models having agrid, it is the grid which both controls the path of the electron beamand focuses the beam.

The transmission of an electron beam from cathode to anode is influencedby electron space charge forces which tend to become significant inconventional x-ray machines at currents exceeding 1 A. In suchconventional machines, the beam is focused on the anode to a spotdiameter ranging anywhere from 0.3 to 2.5 millimeters (mm). In manyapplications, most of the energy from the electron beam is convertedinto heat at the anode. To accommodate such heating, high power medicalx-ray sources often utilize liquid cooling and a rapidly rotating anode,thereby establishing an increased effective target area, permitting asmall focal spot while minimizing the effects of localized heating. Toachieve good thermal conductivity and effective heat dissipation, theanode typically is fabricated from copper. In addition, the area of theanode onto which an electron beam is incident requires a material ofhigh atomic number for efficient x-ray generation. To meet therequirements of thermal conductivity, effective heat dissipation, andefficient x-ray generation, a tungsten alloy typically is embedded inthe copper.

In use, the total exposure from an x-ray source is directly proportionalto the time integral of the electron beam. During relatively longexposures (e.g. lasting 1 to 3 seconds), the anode temperature may risesufficiently to cause it to glow brightly, accompanied by localizedsurface melting and pitting which degrades the radiation output.However, thermal vaporization of the tube's coiled cathode filament ismost frequently responsible for conventional tube failure.

While the efficiency of x-ray generation is independent of the electronbeam current, it is highly dependent on the acceleration voltage. Below60 kV, only a few tenths of one percent of the kinetic energy from anelectron is converted to x-rays, whereas at 20 MV that conversion factorrises to 70 percent. An emitted x-ray spectrum is composed in part ofdiscrete energies characteristic of transitions between bound electronenergy levels of the target element. The spectrum also includes an x-rayenergy continuum, known as bremsstrahlung, which is caused byacceleration of the beam electrons as they pass near target nuclei. Themaximum energy of an x-ray cannot exceed the peak energy of an electronin the beam. Further, the peak of the bremsstrahlung emission curveoccurs at approximately one-third the electron energy.

Increasing the electron current results in a directly proportionalincrease in x-ray emission at all energies. However, a change in beamvoltage results in a total x-ray output variation approximately equal tothe square of the voltage, with a corresponding shift in peak x-rayphoton energy. The efficiency of bremsstrahlung radiation productionincreases with the atomic number of the target element. The peak outputin the bremsstrahlung curve and the characteristic spectral lines shiftto higher energies as the atomic number of the target increases.Although tungsten (Z=74) is the most common target material used inmodem tubes, gold (Z=79) and molybdenum (Z=42) are used in somespecialty tubes.

X-rays interact in several ways with matter. For biological samples, thefollowing two types of interactions are most important: Comptonscattering of moderate-energy x-rays with outer shell electrons; and,photoionizing interactions of inner shell electrons. In these processes,the probability of atom ionization decreases with increasing photonenergy in both soft tissue and bone. For the photoelectric effect, thisrelationship follows an inverse third-power law.

One disadvantage of present x-ray devices used for therapy is the highvoltage required when directed to soft tissue within or beneath bone.One example is in directing x-rays to areas of the human brain, which issurrounded by bone. High energy x-rays are required to penetrate thebone, but often damage the skin and brain tissue. Another example inradiation therapy is in directing the x-rays to soft tissue locatedwithin the body cavity, couched among other soft tissue, or within aninternal calciferous structure. Present high-voltage machines arelimited in their ability to selectively provide desired x-ray radiationto such areas.

Another disadvantage of the high voltage output of present x-ray sourcesis the damage caused to skin external to the affected organ or tissue.Therefore, high voltage devices of present systems often causesignificant damage not only to the target region or tissue, but also toall surrounding tissue and surface skin, particularly when used forhuman tumor therapy. However, since present devices apply x-rayradiation to target regions internal to a patient from a source externalto the target region, such incidental tissue damage is practicallyunavoidable.

Specifically as to brain tissue, which lacks any substantialregenerative ability, the treatment of brain tumors requires precisetechniques to bring about specific tissue destruction. The use ofconventional x-ray devices in brain tumor therapy often lacks theprecision needed in volumetric irradiation, resulting in the damage ofnon-cancerous tissue of the brain and associated glandular structures.

An alternative form of tumor therapy, called brachytherapy, involvesimplanting encapsulated radioisotopes in or near the tumor to betreated. While such use of radioisotopes may be effective in treatingcertain types of tumors, introduction of the isotopes requires invasiveprocedures which have potential side-effects, such as the possibility ofinfection. Moreover, brain swelling may occur in some applicationsbecause the emission from the isotope cannot be controlled. Further,there is no ability to provide selective control of time dosage orradiation intensity. Handling and disposal of such radioisotopesinvolves hazards to both the individual handler and the environment.

Invasive techniques of the brain require precise control of irradiationthrough the choice and concentration of isotopes used. Intracranialpenetration poses a significant risk as is well known in the art.

In view of the above requirements and limitations to the use of x-raysfrom present machines in therapeutic, diagnostic, palliative, orevaluative environments, there remains a need for a relatively small,easily manipulated, controllable, low-energy, x-ray device where thex-ray source can be positioned in proximity to the environment to beirradiated. Such a device operating at low energy and power will besuitable for many of the applications described herein.

Thus, it is an object of the present invention to provide an easilymanipulated, low-power x-ray device.

It is another object of the invention to provide a relatively small,low-power x-ray device having a controllable, or programmable, powersupply.

It is another object of the invention to provide a relatively small,low-power x-ray device which is implantable into a patient for directlyirradiating a desired region of tissue with x-rays.

It is another object of the invention to provide a low-power x-raydevice for irradiating a volume to establish an absorption profiledefined by predetermined isodose contours in order to reduce tissuedamage outside the desired irradiation region.

It is yet another object of the invention to provide a relatively small,surface-mountable, low-power x-ray device for affecting a desiredsurface region with x-rays.

It is yet another object of the invention to provide a relatively small,low-power x-ray device which is partially implantable into a patient fordirectly irradiating a specified region with x-rays.

It is yet another object of the invention to provide a small, low-powerx-ray device and reference frame assembly for controllably positioningan x-ray source within a patient's skull in order to irradiate andtherefore treat a brain tumor.

It is yet another object of the invention to provide a small, low powerx-ray device which can be threaded through existing, irregularly shapedpassageways.

It is yet another object of the invention to provide a small, low powerx-ray device which includes an improved mechanism for directing anelectron beam at a target element.

SUMMARY OF THE INVENTION

Briefly, the invention is an easily manipulated apparatus having alow-power, electron beam (e-beam) activated x-ray source of preselected,or adjustable, duration, effective energy and intensity. In medicalapplications, the apparatus (or "probe") may be fully or partiallyimplanted into, or surface-mounted onto a desired area of a patient toirradiate a region with x-rays. Additionally, the apparatus can beassembled with a variable-thickness x-ray shield to allow irradiationof, and consequent absorption in, a preselected volume, defined by a setof isodose contours, so as to reduce the destructive effects of x-raysoutside the desired irradiation region. The apparatus can be assembledin combination with a reference frame, for example, a stereotacticframe, and an associated coupler for use in the treatment of braintumors.

The apparatus operates at a relatively low voltage, for example, in therange of approximately 10 kV to 90 kV, with small electron currents, forexample, in the range of from approximately 1 nA to 100 μA. To achieve adesired radiation pattern over a desired region, while minimallyirradiating other regions, x-rays are emitted from a nominal, oreffective "point" source located within or adjacent to the desiredregion-to-be-irradiated. In some applications, a low dose rate of x-raysirradiates any part of the desired region, either continually orperiodically, over extended periods of time. For use with a referenceframe for treatment of brain tumors, a high dose rate for single doseirradiation is generally preferred. With the use of a "repeatlocalizer," the single dose can be replaced, if desired, by a series ofhigh dose rate, i.e., fractionated, treatments.

The apparatus includes a controllable, or programmable, power supplylocated outside the desired region-to-be-irradiated to enable variationsin voltage, current, and timing of an electron beam. The electron beamis controlled to pass along a desired beam axis and to be incident on atarget which is preferably located in the patient's body. The axis maybe straight, or curved. The composition and/or geometry of the target,or x-ray emitting, material is selected to provide a customized patternof x-rays. Shielding at the emission site, or around the target, furtherenables control of the energy and spatial profile of the x-ray emissionto match the preselected distribution of radiation throughout thedesired region. A stable and reproducible source of x-rays can becreated with the electron spot either larger or smaller than the target,although the former results in an inefficient use of electrons and thelatter may compromise the spherical isotropy of the emitted radiation.

The present invention further provides a method of treating malignantcells, such as found in tumors, in vivo, utilizing the apparatusdescribed above. Generally, the method involves identifying and locatingmalignant cells with a device generally available in the art, such as bycomputed tomography (CT) scanning or magnetic resonance imaging (MRI). Aneedle-type biopsy of the tumor may be performed to confirm thediagnosis. Then the region of treatment is selected and the radiationdosage determined. Such radiation treatment planning involves definingthe size and shape of the tumor determining precisely its location inthe body, identifying radiation-sensitive critical biological structuressurrounding the tumor, deciding on the proper radiation dosedistribution in the tumor and surrounding tissue and the entry path into the tumor of the implanted portions of the apparatus. For sphericaltumors, treatment planning can be performed manually using CT or MRIdata. However, for more complex geometries, close-by criticalstructures, or higher precision procedures, computer-based "3-D" imageryis preferred. In that case, tumors and critical structures are, forexample, manually or semiautomatically segmented on a series ofdigitized CT scans, and a 3-D composite is rendered, which allowsviewing the tumor from any direction. Various software systems have beendeveloped for radiosurgical procedures, such as those using the linacand gamma knife, and some are commercially available. For example,Radionics Software Applications of Brookline, Massachusetts offers forsale software which images the CRW and BRW stereotactic frame affixed toa graphically transparent skull. Isodose profiles are overlaid on thetumor and other brain tissue. Similar software may be used with theinvention disclosed in U.S. patent application Ser. No. 955,494 whicheffects imaging with respect to a stereotactic frame, for use with thex-ray-radiating electron beam target imbedded in the tumor. Isodosecontours around the target are superimposed on the tumor and adjacenttissue. The absolute radiation dosage delivered along each contour isdetermined by experimental dosimetry performed to calibrate the probe.In these tests, the dose is measured at multiple locations around thetarget immersed in a water tank. Soft tissue is adequately simulated bywater. The dose is measured by an ionization chamber, such as ismanufactured by PTW of Freiburg, Germany, wherein x-ray generated ionscreate a small current which is detected by an electrometer, such as iscommercially available from Keithley Radiation Division in Cleveland,Ohio. Alternatively, the target can be immersed in a biological tissuesimulating phantom. Such plastic, "solid water," phantoms arecommercially available (RMI, Middleton, Wis.) and simulate various bodytissues, e.g., soft tissue of the brain. Either thermoluminescentdetectors (TLD) or calibrated x-ray sensitive film (e.g., gafchromicfilm from Far West Technologies, Goleta, Calif.) can be positioned inthe solid water to measure the dose directly. Using the imaging anddosimetry results from the radiation treatment planning, a low-powerelectron beam source and a selectively shaped x-ray radiation patterngenerating target and shield assembly are positioned within or proximalto a region containing the cells to-be-irradiated, generally tumorcells, for example, in conjunction with a stereotactic frame assembly,such as disclosed in U.S. patent application Ser. No. 955,494. Otherpositioning assemblies, or methods, may be used.

Pursuant to the present invention, the target and shield assemblygeometry and materials are shaped and selected in accordance with thecharacteristics of the desired region-to-be-irradiated. A programmablepower supply is provided, which may be used to vary the voltage,current, and duration of the electron beam source to establish, inaccordance with dosimetry information, a desired electron beam which isdirected to the desired region-to-be-irradiated. Finally, x-rays emittedfrom the target and modified by the shield assembly are transmitted tothe desired region-to-be-irradiated for selective destruction of thecells in that region. By use of a method of signal feed-back, in whichthe x-rays emitted from the target in an backward direction along thepath of the electron beam are monitored by a detector positioned behindthe electron emitter, adjustments in the deflection of the electron beamcan be made to automatically control and optimally position the electronspot on the target.

In particular, the treatment of a brain tumor can be carried oututilizing an apparatus of the present invention comprising thecombination of a low-power x-ray source for generation of a controllableirradiation pattern, with a device for accurately positioning the x-raysource in the brain. The x-ray source can thus be precisely located nearor in the tumor.

The x-ray source, together with the target and shield assembly, of thepresent invention may be used in various body locations to generatecustom-designed irradiation fields for treatment of a variety of typesof tumors. Also, irradiation fields can be customized for each humantreated. However, geometrical similarities for many tumors will allowthis treatment with a standard set of shields.

According to a further embodiment of the invention, the probe can beflexible in nature to allow it to be threaded down existing passagewaysor around obstacles. According one such embodiment, a photoemissiveelement (i.e. a photocathode) is located, along with a target element,in the target assembly. Additionally, a flexible fiber optical cable,which couples light from a laser source to the photocathode, can formthe basis for the flexible probe.

One terminal of a high voltage power supply is coupled to thephotocathode, via an electrical conductor embedded in the fiber opticalcable. The other terminal of the power supply is coupled to the targetelement, via an electrically conductive, flexible, outer sheath formedaround on the fiber optical cable. In this way, an electrical field isestablished which acts to accelerate electrons emitted from thephotocathode toward the target element. As in previously discussedembodiments, the target element emits x-rays in response to incidentelectrons from the photocathode.

BRIEF DESCRIPTION OF DRAWINGS

The foregoing and other objects of this invention, the various featuresthereof, as well as the invention itself, may be more fully understoodfrom the following description, when read together with the accompanyingdrawings in which:

FIG. 1 is a perspective view of a low power x-ray source embodying thepresent invention;

FIG. 2 is a schematic representation of a sheath adapted for use withthe apparatus of FIG. 1;

FIGS. 3A and 3B are a perspective view and sectional view, respectively,of a surface-mountable apparatus embodying the present invention;

FIG. 4 is a schematic block diagram of the embodiment of FIG. 1;

FIGS. 5A and 5B are graphical representations of the x-ray emissionspectrum of tungsten- and molybdenum-targets, respectively;

FIG. 6 is a detailed block diagram of a representative power supply ofthe embodiment of FIG. 1;

FIG. 7 is a detailed schematic diagram of power supply of FIG. 6;

FIG. 8 is a perspective view of a beam steering assembly embodying thepresent invention;

FIG. 8A is a cross-section view of the assembly of FIG. 8, taken alonglines 8a;

FIG. 9 is a perspective view of a brain tumor x-ray treatment systemincorporating a stereotactic frame for positioning the x-ray source;

FIG. 10 is an exploded perspective view of an x-ray source and thecoupling assembly of the system of FIG. 9;

FIG. 11 is a schematic diagram of a representative high voltage powersupply of the x-ray source of FIG. 10;

FIG. 12 is a cross-sectional view of the end of a probe having analternate target assembly which includes an x-ray shield and x-raytarget for producing a stable and reproducible source of x-rays;

FIG. 13 is a cross-sectional fragmentary view of one geometric form ofan x-ray target;

FIG. 14 is a block diagram of a laser milling system for generatingvariable thickness x-ray shields;

FIGS. 15A and 15B are perspective views of a probe and target assemblyfor accurate angular alignment of an x-ray shield;

FIGS. 16 is a cross-sectional view of a low power x-ray source having aninternal beam steering assembly which includes a feedback loop forelectron beam positioning;

FIG. 17 is a cross-sectional view of a low power x-ray source having anexternal beam steering assembly which includes a feedback loop forelectron beam positioning;

FIG. 18 is a cross-section view of the assembly of FIG. 17, taken alonglines 16C;

FIG. 19 is a cross-section view of a mechanical probe positioner forbroad-area irradiation;

FIGS. 20A and B are cross-sectional views of a flexible probe whichincorporates a photoemitter located within the target assembly;

FIGS. 21A-21F show examples of various isodose contours that can beachieved with the invention; and

FIG. 22 is a cross-section view of the end of a probe having analternate target assembly which includes a high impedance shield inclose proximity to the photocathode.

Like numbered elements in each FIGURE represent the same or similarelements.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

The present invention is a relatively small, electron-beam activated,low power x-ray apparatus. The apparatus may be used for medicalpurposes, for example, therapeutic or palliative radiation treatment oftumors, or for other purposes.

With particular regard to medical uses, the apparatus may be fullyimplanted or partially inserted into a preselected internal region of apatient to provide x-ray radiation over selected exposure times.Alternately, the apparatus may be mounted on a surface of a patientexternal to a region to be irradiated. Also disclosed is a method fortreating tumors in a patient, using the apparatus of the invention.

Generally, the apparatus of the present invention includes anelectron-beam (e-beam) activated x-ray source which operates atrelatively low voltages, i.e. in the range of approximately 10 kV to 90kV, and relatively small electron beam currents, i.e. in the range ofapproximately 1 nA to 100 μA. At those operating voltages and currents,the x-ray output is relatively low, and the apparatus may be made quitesmall and be adapted for implantation in medical therapeuticapplications. In view of the low level x-ray output, adequate tissuepenetration and cumulative dosage may be attained by locating the x-raysource adjacent to or within the region to be irradiated. Thus, thex-rays are emitted from a well-defined, small source located within oradjacent to the region to be irradiated. In one embodiment, a low doserate of x-rays may be applied to any part of a tumor, either continuallyor periodically, over extended periods of time, e.g., up to one month.In use with a stereotactic frame for the treatment of brain tumors, ahigher dose rate may be applied to a tumor for shorter periods of time(i.e., on the order of 5 minutes to 3 hours).

The present invention provides interstitial radiotherapy similar to thatachieved with implanted capsules, needles, tubes, and threads containingnatural or artificial radioactive isotopes, known as brachytherapy.However, a programmable power supply may be included in the x-ray sourceof the present apparatus to vary energy, intensity, and duration of theradiation. This differs from brachytherapy in that the intensity andpenetration depth of the x-rays may be changed without surgically orinvasively replacing the isotopes. Furthermore, the present invention isnot limited by the half-life of a particular isotope, and does not posea radiation hazard when turned off.

FIG. 1 shows an x-ray apparatus 10 embodying the present invention.Apparatus 10 includes a housing 12 and an elongated cylindrical probe 14extending from housing 12 along a reference axis 16. The housing 12encloses a high voltage power supply 12A (illustrated in electricalschematic form in FIGS. 6 and 7). The probe 14 is a hollow tube havingan electron beam generator (cathode) 22 adjacent to the high voltagepower supply 12A. Cathode 22 is located in close proximity to an annularfocusing electrode 23 typically at nearly the same potential as thecathode 22. An annular anode 24 is positioned approximately 0.5 cm ormore from the annular focusing electrode 23. A hollow, tubular probe 14extends along the same axis as the cathode, grid, and the hole in theanode. Probe 14 is integral with the housing 12 and extends toward atarget assembly 26. In various embodiments, parts of the probe 14 may beselectively shielded to control the spatial distribution of x-rays. Inaddition, the probe 14 may be magnetically shielded to prevent externalmagnetic fields from deflecting the beam away from the target.

The electron beam generator 22 may include a thermionic emitter (drivenby a floating low voltage power supply) or a photocathode (irradiated byan LED or laser source). The high voltage power supply establishes anacceleration potential difference between the cathode of generator 22and the grounded anode 24 so that an electron beam is established alongthe reference axis 16, through the center hole of the anode and to thetarget assembly 26, with the region between anode 24 and the targetassembly 26 being substantially field free. The beam generation andacceleration components are adapted to establish a thin (e.g. 1 mm orless in diameter) electron beam within the probe 14 along a nominallystraight axis 16.

In a preferred embodiment, the probe 14 is a hollow, evacuated cylindermade of a beryllium (Be) cap and a molybdenum-rhenium, (Mo--Re),molybdenum (Mo) or mu-metal body and a stainless-steel base extension.The cylinder is 16 cm long, with an interior diameter of 2 mm, and anexterior diameter of 3 mm. The target assembly 26 includes an emissionelement consisting of a small beryllium (Be) target element 26A coatedon the side exposed to the incident electron beam with a thin film orlayer 26B of a high-Z element, such as tungsten (W), uranium (U) or gold(Au). By way of example, with electrons accelerated to 30 keV-, a 2.2micron thick tungsten film absorbs substantially all the incidentelectrons, while transmitting approximately 95% of any 30 keV-, 88% ofany 20 keV-, and 83% of any 10 keV- x-rays generated in that layer. Inthe preferred embodiment, the beryllium target element 26a is 0.5 mmthick with the result that 95% of these x-rays generated in directionsnormal and toward the substrate, and having passed through the tungstentarget, are then transmitted through the beryllium substrate and outwardat the distal end of probe 14. While the target element 26A shown inFIG. 3B is in the form of a disc, other shaped elements may be used,such as those having hemispherical or conical outer surfaces.

In some forms of the target, the window element 26A may include amultiple layer film 26B, where the differing layers may have differentemission characteristics. By way of example, the first layer may have anemission (vs. energy) peak at a relatively low energy, and the second(underlying) layer may have an emission (vs. energy) peak at arelatively high energy. With this form of the invention, a low energyelectron beam may be used to generate x-rays in the first layer (toachieve a first radiation characteristic) and high energy electrons maybe used to penetrate through to the underlying layer (to achieve asecond radiation characteristic). As an example, a 0.5 mm wide electronbeam is emitted at the cathode and accelerated to 30 keV- through theanode, with 0.1 eV transverse electron energies, and arrives at thetarget assembly 26 sixteen centimeters downstream from the anode, with abeam diameter of less than 1 mm at the target element 26A. X-rays aregenerated in the target assembly 26 in accordance with preselected beamvoltage, current, and target element 26A composition. The x-rays thusgenerated pass through the beryllium target element 26A in the probewith minimized loss in energy. As an alternative to beryllium, thetarget element 26A may be made of carbon or other suitable materialwhich permits x-rays to pass with a minimum loss of energy. An optimalmaterial for target element 26A is carbon in its diamond form, sincethat material is an excellent heat conductor. Using these parameters,the resultant x-rays have sufficient energy to penetrate into softtissues to a depth of a centimeter or more, the exact depth dependentupon the x-ray energy distribution.

The apparatus of FIG. 1 is particularly adapted for full implantationinto a patient, where the housing 12 has a biocompatible outer surfaceand encloses both a high voltage power supply circuit 12A forestablishing a drive voltage for the beam generator 22, and anassociated battery 12B for driving that circuit 12A. In this case, anassociated controller 12C establishes control of the output voltage ofthe high power supply circuit 12A, in the manner described below.

The apparatus of FIG. 1 may also be used in a manner where only theprobe 14 is inserted into a patient while the housing remains outsidethe patient, i.e., a transcutaneous form. In the latter form, some orall of the various elements shown within housing 12 may alternatively beremotely located.

In the transcutaneous form, the apparatus 10 may be used with anelongated closed end (or cup-shaped) sheath 34, as shown in FIG. 2,having a biocompatible outer surface, for example, fabricated of medicalgrade aliphatic polyurethane, as manufactured under the trademarkTecoflex® by Thermedics, Inc., Woburn, Mass., With this configuration,the probe 14 is first inserted into the sheath 34. The sheath 34 andprobe 14 are then inserted into the patient through the skin.Alternatively, a port may be inserted through the skin and attached toit, as for example a Dermaport® port manufactured by Thermedics Inc.,Woburn, Mass. The probe 14 is then inserted into the port.

The lining of the sheath or port can be configured as an x-ray shield byintroducing barium sulfate or bismuth trioxide, or other x-ray shieldingmaterials, into the sheath. If necessary, the probe 14 and housing 12can be secured to the patient's body to prevent any relative motionduring the extended time of treatment. An exemplary sheath 34 is shownin FIG. 2.

In one embodiment of the apparatus as shown in FIG. 1, the main body ofthe probe 14 can be made of a magnetically shielding material such as amu-metal. Alternatively;, the probe 14 can be made of a non-magneticmetal, preferably having relatively high values for Young's modulus andelastic limit. Examples of such material include molybdenum, rhenium oralloys of these materials. The inner or outer surface of probe 14 canthen be coated with a high permeability magnetic alloy such as permalloy(approximately 80% nickel and 20% iron), to provide magnetic shielding.Alternatively, a thin sleeve of mu-metal can be fitted over, or insideof, the probe 14. The x-ray apparatus 10 can then be used inenvironments in which there are dc and ac magnetic fields due toelectrical power, the field of the earth, or other magnetized bodiesnominally capable of deflecting the electron beam from the probe axis.

In implantable configurations, the power supply 12A and target assembly26 are preferably enclosed in a metal capsule to prevent current flowfrom the x-ray source to the patient. The closed housing 12 and probe 14are, thus, encapsulated in a continuous outer shell of appropriateshielding material such as those mentioned previously.

The high voltage power supply 12A in each of the illustrated embodimentspreferably satisfies three criteria: 1) small in size; 2) highefficiency to enable the use of battery power; and 3) independentlyvariable x-ray tube voltage and current to enable the unit to beprogrammed for specific applications. A high-frequency, switch-modepower converter is used to meet these requirements. The most appropriatetopology for generating low power and high voltage is a flyback voltageconverter working in conjunction with a high voltage,Cockroft-Walton-type multiplier. Low-power dissipation, switch-modepower-supply controller-integrated circuits (IC) are currently availablefor controlling such topologies with few ancillary components.

In order to provide active control of the x-rays, a preferred embodimentof the present invention establishes independent control of cathodevoltage and current without the use of a grid electrode. In that form ofthe invention, an radio frequency ohmic heating current is provided to athermionic cathode 22, preferably using a transformer-coupled 0.6 volt,0-300 mA filament power supply floating at the cathode potential of 40kV.

FIGS. 3A and 3B show an alternative embodiment 10' of the inventionadapted for superficial usage, that is for direct placement on the skinof a patient. This form of the invention is particularly useful forx-ray treatment of skin lesions or tumors, or other dermatologicalapplications. In FIGS. 3A and 3B, elements that correspond to elementsin the embodiment of FIG. 1 are denoted with the same referencedesignations. Apparatus 10' generates an electron beam in a channel 40enclosed within housing 12, where that channel 40 corresponds to probe14. In the present embodiment, of FIGS. 3A and 3B, the target assembly26 (elements 26A and 26B) functions as the anode as well as an x-rayemitter. Otherwise, the apparatus 10' is similar to apparatus 10. Aswith the configuration of FIGS. 3A and 3B, low power x-rays may bedirected to a desired skin region of a patient.

In all of the above-described embodiments, the x-ray emission element ofthe target assembly is adapted to be adjacent to or within the region tobe irradiated. The proximity of the emission element to the targetedregion, e.g. the tumor, eliminates the need for the high voltages ofpresently used machines, to achieve satisfactory x-ray penetrationthrough the body wall to the tumor site. The low voltage alsoconcentrates the radiation in the targeted tumor, and limits the damageto surrounding tissue and surface skin at the point of penetration. Forexample, the delivery of 4000 rads, as is required after a mastectomy,with a 40 kV, 20 μA electron beam, may require approximately 1 to 3hours of radiation. However, since the x-ray source is, in thispreferred embodiment, insertable proximate to, or into, theregion-to-be-irradiated risk of incidental radiation exposure to otherparts of the patient's body is significantly reduced.

Further, specificity in treating tumors may be achieved by tailoring thetarget and shield geometry and material at the emission site. Thistailoring facilitates the control of energy and the spatial profile ofthe x-ray emission to ensure more homogenous distribution of theradiation throughout the targeted tumor.

FIG. 4 is a schematic representation of the x-ray source apparatus 10shown in FIG. 1. In that preferred configuration, the housing 12 isdivided into a first portion 12' and a second portion 12". Enclosedwithin the first housing portion 12' is a rechargeable battery 12B, arecharge network 12D for the battery 12B, which is adapted for use withan external charger 50, and a telemetry network 12E, adapted to beresponsive to an external telemetry device 52 to function in the mannerdescribed below. That portion 12' is coupled by cables to the secondhousing portion 12". The second housing portion 12" includes the highvoltage power supply 12A, controller 12C and the probe 14, as well asthe electron beam generating element 22. In one embodiment, the electronbeam generator includes a thermionic emitter 22 driven by the powersupply 12A. In operation, power supply 12A heats the thermionic emitter22, which in turn generates electrons which are then accelerated towardthe anode 24. The anode 24 attracts the electrons, but passes themthrough its central aperture toward the target assembly 26. Thecontroller 12C controls the power supply 12A to dynamically adjust thecathode voltage, the electron beam current, and temporal parameters, orto provide pre-selected voltage, beam current, and temporal parameters.

Also illustrated, is an alternative electron beam generator whichincludes a photoemitter 22 irradiated by a light source 56, such as adiode laser or LED, powered by a driver 55. The light is focused on thephotoemitter 22 by a focusing lens 58.

In the illustrated embodiment, device 52 and network 12E cooperate topermit external control (dynamic or predetermined) control over thepower supply 12A and temporal parameters. In embodiments when thehousing 12" is not implanted, but where only probe 14 extends into apatient's body, the controller 12C may directly be used to controloperation; in that case there is no need for network 12E.

In an important aspect of the invention, the target assembly 26 may beshaped to emit x-rays in a radiation pattern in a predetermined spectralrange, and having a predetermined spatial distribution. This spectraltarget shaping may be achieved in part by selecting target materials ofknown characteristics. For example, as shown in FIGS. 5A and 5B, theemission spectrums for tungsten targets (FIG. 5A) and molybdenum targets(FIG. 5B) are distinct. FIG. 5A shows the x-ray emission spectrum from atungsten target tube operating at 30 and 50 kV. Note that thebremsstrahlung spectrum predominates, and that x-rays are supplied in awide energy range. FIG. 5B shows the emission spectrum from a molybdenumtarget tube, also operating at 30 and 50 kV. Note the near absence ofbremsstrahlung x-rays. Note also that the change in tube potential from30 to 50 kV results in a minor change in the shape of the emissionspectrum from a molybdenum target x-ray tube. Thus, the x-ray spectralemission from target assembly 26 may effectively be shaped by selectingthe target material to provide the desired radiative penetration oftissue, e.g., the tumor.

The x-ray spatial distribution may be also shaped by altering thegeometric configuration of target element 26A. By way of example, thetarget element 26A may be shaped such that the electrons directed fromthe anode will be incident at a predetermined angle or may beselectively directed to different areas of the region from whichemission is to occur. By way of further example, the target element 26Acan be fabricated to be thick enough to be substantially opaque toelectrons but thin enough to be substantially transparent to x-rays.More specifically, if a spherical gold target element having a diameterof 0.5 μm and a 40 kV electron beam is employed, substantially all ofthe electrons are stopped by the target element and substantially all ofthe x-rays generated in the target element can escape.

The x-ray spatial distribution can also be shaped by incorporating anx-ray transmissive shield, having a variable thickness profile, into thetarget assembly 26. FIG. 12 shows a probe 14 having an alternate targetassembly 126, for use with the x-ray apparatus 10 shown in FIG. 1, whichincorporates such a shield. In the illustrative embodiment, the probe 14is substantially similar to the probe 14 shown in FIG. 1, except for thetarget assembly 126. Target assembly 126 includes a probe tip 126A madeof a material (e.g. Be) which is nearly transparent to x-rays, and anx-ray target 126B for generating a source of x-rays upon irradiationwith an electron beam, attached to the probe 14 along a probe axis 16 atthe end distal to the cathode 22 and anode 24 (shown in FIG. 1). In thepreferred form, the outer surface of the probe tip 126A is convex, andpreferably hemispherical, as in the illustrated embodiment, althoughother convex shapes can be used. The target assembly 126 is fabricatedsuch that the outer diameter of the probe tip 126A is less than theouter diameter of the probe 14. A variable thickness x-ray shield (orshadow mask, as it is sometimes referred to in the art) 128 and anunderlying shield carrier 128A are positioned over the probe tip 126A ofthe target assembly 126. At the junction of the target assembly 126 andprobe 14, the outer diameter of the target assembly 126 substantiallymatches that of probe 14.

The x-ray shield 128 is made from a material which is not completelyx-ray transparent (i.e. at least partially x-ray absorptive, such asheavy metals), and is supported by the shield carrier 128A. The x-rayflux from any point of the target assembly 126 is dependent in part uponthe thickness of the x-ray shield 128 along an axis extending from thetarget 126B and passing through that point. Thus, in accordance with theinvention, a selective restriction in thickness of the x-ray shield 128is used to generate spatially-variable x-ray dose distributions.

In a preferred embodiment, the probe 14 has an outer diameter of 3 mmand an inner diameter of 2 mm, and is typically 10 to 16 cm long. Thetarget carrier 126C is made of beryllium and has a hemispherical tip126C' with a radius 0.8 mm, the probe tip 126A is made of beryllium andhas a thickness 0.5 mm. The shield carrier 128A is made of a lightelement, such as beryllium, magnesium, aluminum, or carbon and has athickness 0.2 mm, and the shield 128 has a thickness in the range 0 to0.1 mm if made of gold.

The x-ray target 126B is a small disk (e.g., 0.1 mm diameter) of anx-ray emissive material (e.g., a metal with a high atomic number such asgold) deposited in the center of the target carrier 126C. As will bediscussed in further detail below, the size of the x-ray target 126B maybe small relative to the diameter of the electron beam established alongthe probe axis 16, so that the source of x-rays produced is defined bythe position of the small target and not by the position or size of theelectron beam. This feature permits illumination of the x-ray shield 128with a reproducible and stable source of x-rays. However, for anelectron beam whose spot on the target 126B is larger than the target126B, there is a loss of efficiency in generating x-rays. Such a losscan be avoided by focusing the beam to a small spot comparable to thesize of the target 126B, and controlling its position on the target 126Bby suitable means.

The spatial resolution of the preselected irradiation volume which canbe obtained by rising the shield 128 is limited by several factors,including the penumbra due to the nonzero size of the x-ray source; theinstability of the size and position of the x-ray source due tocorresponding instability in the x-ray generating electron spot; thescattering of the x-ray deposited energy in the irradiated volume; andthe probe-to-probe reproducibility of the x-ray source and its positionrelative to the shield 128.

The penumbra is determined by the ratio of the size of the x-ray sourceto its distance from the shield 128. For a uniform source, a preferablerange for this ratio is on the order of 1/20 to 1/3, depending on thescattering behavior. The stability of the size of the x-ray source andits position is preferably a small fraction of the optimum source todistance ratio.

One method of establishing an acceptable penumbra and registration ofthe shield x-ray source is to control the position and size of the x-raysource by controlling the focus and deflection of the incident electronbeam along axis 16. For instance, the electron beam can be focused to aspot on the x-ray emissive surface of target 126B, the diameter of thefocal spot thus being the size of the x-ray source. This method requiresnot only that the spot size be correct, but that the position of thespot relative to the x-ray shield 128 be accurately known andmaintained.

In this embodiment, the target can theoretically be as large asfabrication convenience dictates. However, in a preferred embodiment,the x-ray target 126B is substantially the same size or only slightlylarger than the electron beam.

In order to ensure that the electron spot position, relative to theshield, is both temporally stable for any given miniature x-ray systemand spatially reproducible in all other systems that are in use,accurately placed fiducial marks can be used together with electron beamdeflectors to locate the electron spot relative to the shield. Such afiducial mark consists of an edge, defining a boundary between tworegions which have very different behavior in an electron beam. Forexample, in the present instance, a boundary between the target material126B, such as Au, and the target carrier material 126C, such as Be, canserve as a fiducial edge. The relevant difference in behavior is that Auis significantly more x-ray transmissive than Be, when exposed to a highenergy electron beam. As the beam passes across the mark, an x-raydetector can sense the difference of x-ray intensity and generate acorresponding control signal for application to the beam deflectors.

The x-ray detector can be embedded in a feedback control loop to servothe beam onto the target and preferably, the center of the target, asviewed from the electron source. In one such configuration, where thetarget position is generally known with respect to the beam path, but isif desired to center the beam path on the target, the beam may first beswept across the target in a first (x) direction which is orthogonal tothe beam path. As the beam passes the fiducial edges of the target (forexample, as the beam encounters the target during the sweep, and then asthe beam leaves the target), the controller identifies the position ofthe fiducial edges and determines an x-component of a control signalrepresentative of the mid-point between the two fiducial edges in thex-direction sweep. Then the beam is positioned in accordance with thatcontrol signal component (i.e. mid-way between the detected x-sweepfiducial edges), and swept in a second (y) direction orthogonal to thex-direction and the beam path. During the y-direction sweep, fiducialedges are detected and a y-component of a control signal is determinedwhich is representative of the mid-point between the two fiducial edgesdetected during the y-direction sweep. The x- and y-components are thenused to control the beam to be centered in the target.

In a case where the target position is not initially known with respectto the beam path, the relative position may be quickly established byraster scanning the beam until the target is encountered in anx-direction sweep, or scan. Then, in response to the detection offiducial edges in that sweep, a mid-point is determined and the beam ispositioned to that mid-point position and then swept in the y-direction,i.e. along the perpendicular bisector of a line connecting the fiducialedges of the identified sweep. In response to detection of fiducialedges in that y-sweep, a y-direction mid-point is determined and controlsignals representative of the x- and y-direction mid-points are used tocenter the beam on the target. Although described above for determiningthe center of a target, other desired reference points on the target maybe determined and the beam deflected to be incident on those points.

Another way to establish proper source position, and hence ensure thespatial resolution of a shielded radiation field for all systems, is touse a small x-ray target 126B which is the size of the desired x-raysource. Although, in principle any size electron spot can be usedwithout degrading the spatial resolution of the shielded radiationfield, it is desirable to make the spot the same size or smaller thanthe target 126B in order to maximize the energy conversion for electronsto x-rays and hence reduce the time to treat patients or to perform anyother desired task using the shielded x-ray source. In this context, ifthe spot size is defined such that 90% of the electrons in the spot arecontained in the so-defined spot size, then making such a spot equal tothe small target size would be optimum in the sense that a smaller spotwould not significantly improve the system efficiency. In practice, itis often the case that smaller spots can only be obtained by decreasingthe current density. In such a case it may not be desirable to make thespot as small as the target. In any event, the use of a small targetensures that all x-ray probes using a shield to define a radiation fieldwill have substantially the same spatial resolution and positionrelative to the probe tip.

As shown in FIG. 12, the target carrier 126C fits snugly into the end ofthe probe tip 126A. In the illustrated embodiment, the x-ray target 126Bis deposited on the target carrier 126C before being inserted into theprobe tip 126A. In instances where the probe tip 126A has been attachedto the body of the probe 14 prior to placement of the x-ray target 126Band target carrier 126C, the target carrier 126C can be fabricated suchthat inner diameter of the probe 14 is slightly greater than the outerdiameter of the target carrier 126C in order to make insertion down thebody of the probe 14 easier.

It is generally desirable that the target carrier 126C fit tightly intothe probe tip 126A in order to ensure mechanical integrity of thestructure. This can be achieved, for instance, by making the parts to"press fit" or by utilizing thermal expansion to clamp the two partstogether. In the latter case a cold target carrier 126C (e.g., cooled byliquid nitrogen) is inserted into a relatively hotter (e.g., roomtemperature) probe tip 126A. As the parts reach thermal equilibrium theyfirmly clamp together.

In an alternative embodiment, the probe tip 126A can be fabricated toinclude an integral target carrier. The probe tip 126A is attached tothe probe 14 subsequent to the placement of the x-ray target 126B.

The x-ray target 126B should be deposited on the target carrier 126Cnormal to the probe axis 16, and at the center of the concentrichemispherical surfaces which define the end of the probe tip 126A. Thisconcentricity of placement of the x-ray target 126B greatly simplifiesthe calculation required to design the variable thickness x-ray shield128 to give desired x-ray isodose contours. As used herein, the termisodose contour refers to a surface of a three-dimensional volume onwhich every point experiences the same x-ray absorption per unit mass oftissue.

Since the x-ray target 126B can be deposited on the target carrier 126Cbefore insertion into the probe 14, any of several methods can be usedto form an x-ray target 126B at the center of the target carrier 126C.One method of fabricating such an x-ray target 126B is to evaporate ahigh-atomic-number metal through a shield which is inserted into thecavity in the target carrier. The shield can consist of a disk with acentral aperture corresponding to the x-ray target 126B and throughwhich the metal is deposited on the target carrier 126C.

In addition to considerations of x-ray source size and position relativeto the x-ray shield 128, it is also necessary to account for x-rayabsorption in the x-ray target 126B itself in a direction tangential tothe plane of the x-ray target 126B. Such absorption can be reduced bymaking the x-ray target 126B a curved surface instead of a flat surface.For example, FIG. 13 shows a quartersphere depression in the targetcarrier 126C which serves to define the form of the x-ray target 126B.The curvature of the x-ray target 126B serves both to reduce theabsorption of x-rays in the target and also to spread out any remainingangular dependence of x-rays emitted from the x-ray target 126B. The netresult can be a much more isotropic emission of x-rays from the x-raytarget 126B which illuminates the x-ray shield 128 located on the shieldcarrier 128A. The curved target shape shown in FIG. 13 is only oneembodiment; other effective shapes may also be used, such as ahemisphere, or a spherical section in combination with a truncated cone.

When the target 126B is deposited in a depression, it can be fabricatedwith the target carrier 126C in situ within the probe tip 126A, or as anintegral part of the probe tip 126A. An evaporative deposition can coatthe depression and surrounding surfaces 126D. The high-atomic-weightmetal deposited on surfaces 126D can, subsequently be removed byscraping the surface with a flat scraper, which does not contact thedepression.

There are applications for the x-ray probe of the present inventionwhich require a broad-area source instead of a point source of x-rays.For example, the resection of a small breast tumor may remove tissue formany centimeters surrounding the focal point of the tumor. Followingresection it may be desired to irradiate the "tumor-bed" in order tokill any remaining tumor cells at the periphery of the resection. In apreferred embodiment, in order to reduce tissue damage beyond thedesired irradiation volume, the broad-area irradiation is carried outwith an x-ray apparatus utilizing an x-ray shield 128 substantiallysimilar to that shown in FIG. 12.

Broad-area radiation can be easily obtained by placing the targetassembly 126 of the probe 14 at a distance from the surface to beirradiated. The solid angle of forward radiation from the targetassembly 126 can be controlled with an x-ray shield 128. The thicknessof the shield 128 at each point is determined so that a substantiallyuniform radiation pattern is obtained. The target assembly 26 can beemployed in a similar fashion.

FIG. 19 shows a mechanical positioner 300 for use with an x-rayapparatus of the present invention to achieve the precision requiredbetween the target assembly 26 or 126 and the irradiated surface(tissue). The mechanical positioner 300 comprises an interface plate 302which contacts the tissue, and is made of some material which istransparent to x-rays, such as Be, C, or plastic. The interface plate302 is attached to the probe 14 by means of an x-ray opaque back plate304. To further pattern a specific radiation field, the surface of thenormally x-ray transparent interface plate 302 can be rendered partiallyx-ray opaque by way of an x-ray shield in a manner similar to the x-rayshield 128 described above.

Another application for such a broad-area x-ray source is intercavityradiation within the body, such as the inside of the bladder. In such acase the interface plate 302 between the tissue and the broad-area x-raysource can be an inflatable balloon, extending down the probe 14 so thatthe target assembly 126 is at the center of the balloon. In this case,there would be no opaque backplate 304.

FIGS. 21A-21F depict examples of various isodose contours that can beachieved with the present invention. Specifically, FIG. 21A shows theprobe 14 adapted to deliver isodose contours which form a sphere ofradiation 300 centered about the probe tip 126. FIG. 21B shows the probe14 adapted to deliver a sphere of radiation 302, wherein the probe tip126 is offset from the center of the sphere 302. FIG. 21C shows theprobe 14 having a tip 126 adapted to deliver a radiation field in theshape of an oblate ellipsoid (i.e., a "pancake" shape), as shown inperspective at 304A and looking along axis 305 at 304B. FIG. 21C depictsthe probe 14 having a tip 126 adapted for delivering a radiation fieldin the shape of a prolate ellipsoid (i.e., a "cigar" shape), as shown inperspective at 306A and along axis 307 at 306B. As shown in FIG. 21D,the probe 14 enters the ellipsoid 306A along its minor axis. FIG. 21Eshows the tip 126 also adapted for delivering a radiation field in theshape of a prolate ellipsoid. The ellipsoid is shown in perspective at308A and along axis 309 at 308B. As can be seen, the probe 14 enters theellipsoid 308A along its major axis. FIG. 21F depicts the probe tip 126adapted for delivering an asymmetric radiation field shown inperspective at 310A and along axis 311 at 310B.

The design of a variable-thickness x-ray shield 128 for generatingx-radiation principally within predetermined isodose contours willgenerally begin with digital data describing the size and shape of thedesired irradiation volume (such as a tumor) which has been obtained bysome method of imaging such as CT scan or Magnetic Resonance Imaging.From such data, and a knowledge of the x-ray absorption properties ofthe probe materials and of the shielding material used, the details ofthe thickness profile of the shield can be calculated. In general theisodose contours can assume many shapes and sizes and need not besymmetrical.

Various methods can be used to translate the design data into a physicalshield. One method would be to use laser milling techniques. Forinstance, a hemispherical shield carrier 128A is coated with a layer ofa metal with a high atomic number (e.g., Au) about 100 μm thick, thethickness of the shielding material deposited on the shield carrier 128Abeing well controlled in order to know how much material to remove in asubsequent milling process. One method of achieving a high degree ofthickness control is to deposit the x-ray absorptive material byelectroplating.

FIG. 14 shows a laser milling system 200 for generating an appropriatevariable thickness x-ray shield 128 for delivery of predetermined x-rayisodose contours. It is well known that intense laser pulses can removesurface layers of metal. The laser milling system 200 of FIG. 14comprises a mechanical positioning apparatus, shown generally asposition controller 202, which systematically presents all of thesurface points of the shield carrier 128A to a laser beam 204. Forinstance, the x-ray shield 128 and shield carrier 128A can be rotatedabout the probe axis 16 or an axis 212 which is normal to the probe axis16. In a preferred embodiment, a microprocessor 210 has direct controlover the motions of the position controller 202, and information as tothe current position of the surface of the x-ray shield 128 istransmitted back to the microprocessor 210 to verify the specifiedposition.

The specifications of the x-ray shield, i.e. the thickness profile, iscalculated prior to the milling process, and from this data, themicroprocessor 210 issues commands to a laser controller 208, whichdrives a laser 214, as to how much power is required to remove thecorrect amount of the shielding material at a particular irradiatedsurface point on the x-ray shield 128.

If the shielding material is entirely metallic, a powerful and expensivelaser may be required in order to complete the milling process in anacceptable length of time. The preferred laser for these conditions isan excimer laser. However, when the shielding material consists of metalparticles suspended in an organic material such as polyimide, then amuch lower power laser, such as a nitrogen laser, may be used.

In another embodiment, the variable thickness x-ray shield 128 can begenerated by controlled vapor deposition of the shielding material. Thistechnique is also amenable to automation, and the pattern of depositioncan be controlled by a microprocessor driven system.

In another embodiment, the shield material is first plated onto thecarrier to the required maximum thickness of about 100 μm for gold, andthen machined with a high accuracy CNC machine tool. This embodiment hasthe advantage of using a simple mechanical process and eliminates theneed for an on-line gauging system as required for laser milling.

FIGS. 15A and 15B show one embodiment of a probe design which allowsaccurate angular alignment of the shield carrier 128A and thus, thex-ray shield 128 with the probe 14. A mechanical key, shown in the formof a tab 140 in the probe 14 and a corresponding groove 142 in thetarget assembly 126, can be provided between the two, to ensure accuratepositioning of the x-ray shield 128 and the probe 14 in order to orientthe x-ray emission pattern with the geometry of the desired irradiationvolume. As one skilled in the art will appreciate, the keyingarrangement of FIGS. 15A and 15B can also be used in combination withthe target assembly 26 of FIG. 1.

As a further feature of the invention, steering may be used to directthe emitted electron beam to selected surfaces on the emission element,for example, where the target has different emission characteristics indifferent spatial regions. Control of the electron beam may be achievedunder the control of telemetry, or by pre-programming the power sourceprior to implantation of all or part of the apparatus 10.

FIG. 8 shows an exemplary electrostatic beam steering assembly. In theillustrated embodiment, the cathode 22 generates electrons in a mannerconsistent with the above-described embodiments. The electrons areaccelerated through a focusing electrode 23 toward the anode 24, andpass through an aperture 24A toward the target assembly 26. Enroute totarget assembly 26, the electrons pass through an electrostaticdeflection assembly 30, shown in cross-section at FIG. 8A. The assemblyincludes four deflectors 32. By varying the voltages applied to theopposing pairs of the deflectors 32, the electrons of the beam enteringthe assembly along axis 16A are deflected, or "steered" as they traveltoward the target assembly 26 along axis 16B. Thus, the beam axis may becontrolled to be straight or curved, as desired. As described below,electromagnetic techniques may alternatively be used to establish beamsteering. In the latter case, the electrostatic deflective plates 32 maybe replaced with magnetic deflector coils which are driven by currentsto establish magnetic fields necessary to achieve a featured beamdeflection.

In another form of the beam-steering embodiment, rather than passthrough an electrostatic deflection assembly 30, the electron beampasses through a set of magnetic field-generating coils. The coils canbe arranged in a configuration similar to the electrostatic deflectionplates of the assembly 30. By varying the current through the coils, theresultant magnetic field is produced in a predetermined manner so as toinfluence the path of the electron beam.

In such a fashion, the electron beam may be steered to hit certainphysical locations on a cone-shaped target assembly (FIG. 8), or atarget of any other specific geometric configuration. By way of example,in the illustrated embodiment, a beam hitting the angled side of targetassembly 26 will result in x-rays emitted off to that side, with littleor no incidental radiation transmitted through to the opposite side ofthe target assembly.

In another form of the beam-steering embodiment, the x-ray emissioncharacteristics may be controlled by spatially varying the emissionparameters (such as radiation peak vs. energy) of the target assembly.By changing the emission peak (as a function of energy) at variouspoints in the target assembly 26, for example, with a "bullseye" spatialpattern, the beam may be steered to regions of relatively high energyx-ray emission, or to regions of relatively low energy x-ray emission.Thus, the beam may be selectively directed to regions of the targetassembly to achieve the required x-ray emission characteristic anddirection.

As one skilled in the art will appreciate, the beam steering assembly ofFIG. 8, can also be used in combination with the target assembly 126 ofFIG. 12.

FIGS. 16, 17 and 18 show an alternate beam steering assembly whichincludes a feedback loop system to accurately position the electron beamon the x-ray target 126B. In the illustrative embodiment, the deflectionassembly 30 is substantially similar to that shown in FIG. 8, (exceptthat it is a magnetic deflection system located outside of the probe)and an x-ray detector 142 is arranged to monitor x-rays emitted from thex-ray target 126B. The x-ray detector 142 can be positioned off axiswith the electron beam, as shown, or placed on axis behind the cathode22.

Changes in the trajectory of the electron beam can be measured whenthere are concomitant changes in the x-ray emission from the target126B. A deflection controller 144, which is preferably driven by amicroprocessor, can utilize the data from the x-ray detector 142 and, bycontrolling the voltages applied to the deflectors 32 of the deflectionassembly 30, can appropriately position the electron beam.

For instance, the feedback loop system can be used to center theelectron beam on a small x-ray target 126B. However, while a change inthe monitor signal does indicate that the center of the beam has movedfrom the center of the target, there is no immediate information as towhich direction the movement has taken place. Hence it may be necessaryto periodically deflect the beam in a known direction and observe thebehavior of the monitor signal in order to recenter the beam.

The monitor signal required to keep the beam positioned on the x-raytarget 126B can be obtained by placing an x-ray detector 142 behind theelectron optics 138 to monitor x-rays which are emitted back along theaxis 16 of the probe 14. In FIGS. 16 and 17, the monitored x-rays 140are shown to pass to one side of the electron optics 138. However, ifthe cathode is thin enough to be transparent to x-rays, it is possibleto design the system such that the x-rays 140 pass through the electronoptics 138 and the cathode 22. The detector 142 can be placed eitherwithin or outside of the housing 12 as shown in FIGS. 16 and 17,respectively. As illustrated in FIG. 17, if the detector 142 is locatedoutside of the housing 12, an x-ray transmissive window 148 should belocated in the wall of the housing to provide optical coupling of thedetector 142 and x-ray target 126B.

After the beam has been accurately centered on the x-ray target 126B,the feedback system of FIGS. 16 and 17 can be used to optimize theelectron-beam focus for maximum output of x-rays. For instance, this canbe accomplished by maximizing the signal monitored by the feedbacksystem by using the deflection controller 144 to adjust the voltages onthe focus elements (such as focusing electrode 23) of the electronoptics 138.

The feedback system illustrated in FIGS. 16 and 17 can also be used withthe target assembly 26 shown in FIGS. 1 or 8. By way of example, thefeedback systems can be used to position the electron beam so as to beincident upon a particular point of an emission element having regionsof different emission characteristics (such as the bullseye spatialpattern described above). Additionally, the feedback system can beemployed to control the acceleration voltage of the electron optics.

As shown in the above-described embodiments, the apparatus 10 of FIG. 1includes a power supply 12A. FIG. 6 is a block diagram of arepresentative supply 12A. FIG. 7 shows a more detailed schematic of thesupply of FIG. 7. As shown in FIGS. 6 and 7, that embodiment includes aflyback switching converter and regulator 280, a 30:1 voltagetransformer 282 coupled to a control voltage (or high voltage multiplierinput) terminal 282A and a 10 stage voltage multiplier 284 coupled to ahigh voltage terminal 22A, and adapted to drive the filament of athermionic emitter 22. A filament radio frequency power driver andvoltage-to-frequency (V/F) converter 290 and an associated radiofrequency filament driver 292 are coupled through current controlterminal 292a and capacitor C_(o) by way of a filament drive circuit 286to the filament of emitter 22.

A difference amplifier 294 establishes a current feedback loop bydriving the radio frequency power driver and V/F converter 290 inresponse to the detected difference between a current feedback signal online 295 and an applied emission control signal on line 296. The lattersignal may be selectively controlled to establish a desired temporalvariation in the x-ray tube cathode current in filament of emitter(thermionic cathode) 22.

A high voltage amplitude feedback loop is established by the switchingconverter and regulator 280 in response to the detected differencebetween a voltage feedback signal on line 297 and an applied highvoltage control signal on line 298. The latter signal may be selectivelycontrolled to establish a desired amplitude variation of the potentialat the filament of emitter (thermionic cathode) 22.

A more detailed description of the power supply shown in FIGS. 6 and 7is provided in U.S. Pat. No. 5,153,900 and also in parent applicationU.S. Ser. No. 955,494.

FIG. 9 shows an exemplary system 300 adapted for x-ray treatment ofbrain tumors. System 300 includes a stereotactic frame 302 incombination with a low-power x-ray device 10A coupled thereto. In thatconfiguration, x-ray device 10A is generally similar to the x-ray device10 shown in FIG. 1, but has a cylindrical geometry. Correspondingelements of the two x-ray devices 10 and 10A are identified with thesame reference designations. In general, stereotactic frames provide afixed reference structure relative to the cranium of a patient. Whilethe preferred embodiment described above is particularly adapted for usewith this stereotactic frame, other embodiments of the invention mightbe similarly adapted for use with this or other frames or with generalreference frames, for example, one establishing an operating fixturefixedly referenced to a part of the body other than the head. In theillustrated embodiment of FIG. 9, the stereotactic frame 302 issubstantially similar to the Cosman-Roberts-Wells system manufactured byRadionics Inc., Burlington, Mass.

In the illustrated embodiment, the frame 302 establishes a reference XYZcoordinate system disposed about a desired origin point 0. The frame 302includes a generally U-shaped support element 304 defining a referenceplane. Four arms 306A, 306B 306C and 306D (not shown) extend out fromsupport frame 304. Each arm has a positioning pin 308. The pins 308extend generally towards each other from the respective distal tips ofarms 306A, 306B, 306C and 306D. In use, the four pins 308 are positionedagainst a patient's skull to establish a fixed positional relationshipbetween the frame 302 and the patient's cranium. Thus, the frame 302defines the reference XYZ coordinate system with respect to thepatient's cranium.

An x-ray device support member 310 is coupled to the support element 304by way of a pair of rotational coupling assemblies 312 and a pair oflinear coupling assemblies 314. The x-ray device support member 310includes an arcuate support track 310A. An x-ray device 10 is coupled tosupport track 310A by a coupling assembly 316. Coupling assembly 316provides controlled movement of the x-ray device 10 on a circular pathalong track 310A and between an inner limit point and an outer limitpoint along axes (exemplified by axis 316') extending radially inwardfrom the circular path of arcuate track 310A toward the origin point O.

In addition, rotation about the hubs of rotational coupling assemblies312 allows the x-ray device support member 310 to be rotatably movedabout the X axis. The x-ray device support member 310 is translocatablein a direction normal to the plane defined by the X and Y axes (the X-Yplane) by movement along tracks 314A, of the linear coupling assemblies314. In the illustrative embodiment, a T-groove in tracks 314A mateswith a tenon of block 314B which is affixed to member 304, permittinglinear motion in the direction perpendicular to the X-Y plane. Setscrews 332 in block 314B may be adjusted to lock the x-ray devicesupport member 310 at a set height relative to the support frame 304.

X-ray support member 310 may be moved in the direction of the Z axis bymovement of the tenons extending from member 310 in tracks 304A ofsupport element 304. A controlled position of the member 310 along thetracks 304A can be established using locking screws 334.

In addition, support element 304 can be adjustably positioned in thedirection of the X axis by sliding member 304 relative to its supportmember 305, and may be adjustably positioned with three degrees offreedom to establish a desired location of origin point O within theskull of a patient.

The coupling assembly 316 is shown together with an x-ray device 10A, inexploded form, in FIG. 10. As shown, the coupling assembly 316 includesa receiver block 316A, a bushing element 316B, together withcomplementary shaped portions of the x-ray device 10A. As shown, thecentral axis 16 of probe 14 of x-ray device 10A is coaxial with axis316'. The electron beam of probe 14 is nominally coaxially with axis316', but may be adjustably varied as described above in conjunctionwith FIGS. 8, 8A, 16, 17, and 18 and below in conjunction with FIG. 10.

The cylindrical bushing element 316B is positioned partially within andcoaxially with the receiver block 316A. The bushing element 316B isslidable (in the direction of radial axis 316') and may be selectivelylocked in place relative to block 316A using a set screw 318A. Thebushing element 316B includes a central bore (with diameter D) extendingalong its central axis.

As noted above, the x-ray device 10A is similar to the x-ray device 10shown in FIG. 1, but has a generally cylindrically shaped housing 12;the probe 14 includes a cylindrical shoulder portion 14A (having adiameter slightly less than D) immediately adjacent to housing 12, witha main portion with a small diameter (3.0 mm in the preferredembodiment). With this configuration, the x-ray device 10A may bepositioned with its axis 16 coaxial with axis 316' and the shoulderportion 14A slidingly positioned within the bore of bushing element316B. The relative position of x-ray device 10A may be fixed along axis316' using set screws 320 of element 316B.

X-ray device 10A may include a magnetic deflection subsystem for itselectron beam. The deflection subsystem includes magnetic deflectioncoils 32 as shown in FIG. 18 positioned about axis 16 within shoulderportion 14A. These coils are driven to adjustably control the positionof the beam axis so that the beam is incident on the target of assembly126 (shown, for example in FIGS. 16 and 17) in a desired manner. In thepreferred form, radiation generated by device 10A is monitored (forexample, by x-ray detector 142 shown in FIGS. 16 and 17, and/or an x-raydetector positioned outside the patient) and the deflector coils aredriven accordingly by steering control currents on deflection X1, X2, Y1and Y2 lines applied to the deflection coils, shown in FIG. 11.

As shown in FIG. 9, the microprocessor-based controller may not bedisposed within the housing 12, but located external to the housing 12in a control unit 342. Control unit 342 is coupled to x-ray device 10Aby way of cable 342'. The elongated probe 14 of x-ray device 10 isconfigured so as to allow the probe 14 to pass through the track left bya biopsy needle, thereby permitting easy insertion of the probe 14 intothe brain of a patient. For tumors composed of hard tissue, and where abiopsy needle smaller in width than the probe 14 is used, properpenetration into the tumor may require first widening the track left bythe biopsy needle with intermediate sized needles.

With this configuration, the tip of probe 14 contains the x-ray emittingtarget and can be moved in and out relative to the cranial insertionsite by movement along the axis 316'. The x-ray device 10A can besecured at a given position along by set screws 318A and 320. The lengthof probe 14 of x-ray device 10A is chosen such that the tip of probe 14,when fully inserted down to the lower limit point along the axis 316' of316A, exactly contacts the origin point O; when the x-ray apparatus 10is fully withdrawn to the upper limit point along axis 316', the distaltip of the probe 14 is intended to be outside the patient's skull. Thecoordinates of the arcuate support track 310A can be set such that theorigin point O is located at the desired epicenter of irradiation. Thus,by the rotation of x-ray device 10A support member 310 and thepositioning of the x-ray device 10A along the circumferential track ofthe arcuate support track 310A and along axis 316', a user can choosethe appropriate path (preferably of least destruction) for insertion ofprobe 14 into a patient's skull, the tip of probe 14 always contactingthe origin point O upon full insertion of the probe 14 to the lowerlimit point.

FIG. 11 shows a schematic diagram of a preferred high voltage powersupply 12A for use with the x-ray device 10A of FIGS. 9 and 10. In thatpower supply, the HV drive signal is a 0 to 9 Volt pulse densitymodulated drive signal. This signal drives the Flyback Switching FieldEffect Transistor (FET) Q1, which in turn drives the HV Flybacktransformer. The HV Flyback transformer steps up the +12 Volts toseveral thousand volts. The HV multiplier, D1 to D28, in turn steps upthe voltage to the desired output voltage of 15 to 40 kV. The voltagefeedback line provides feedback information to controller 12C, so thatthe output voltage of the HV multiplier can be held at a constant value.

The Filament + and - lines provide complementary 9 Volt 250 kHz squarewave drive signals to FET's Q2 and Q3. These FET's chop the variableFilament DC voltage into an AC voltage, and drive the Filament/HVIsolation Transformer T2. Using a high frequency signal to drive thistransformer permits a single turn secondary to drive the x-ray tubefilament. This in turn permits miniaturizing the transformer whilemaintaining the necessary high voltage isolation. The current FB lineallows controller 12C to sense the beam current, and the controller thenadjusts the Filament DC Voltage for the desired beam current, byproviding the appropriate heating current to the thermionic emitter 22.The Deflection X1, X2, Y1, Y2 lines provide current drive signals to themagnetic beam deflection coils.

As discussed above with respect to FIG. 1, the apparatus 10 includesbeam generation and acceleration components to generate and accelerateelectrons, prior to those electrons entering the probe 14. The generatedelectron beam then flows through probe 14, impacts the target 26b, andthereby produces x-rays. In the absence of magnetic fields, theelectrons flowing through the probe 14 follow a straight-linetrajectory. Consequently, the probe 14 is typically rigid without anybends.

However, in certain medical applications it is beneficial to use aflexible probe. One such application involves threading the x-ray sourcedown an existing pathway, such as the trachea. Another such applicationinvolves maneuvering the x-ray source around critical structures, suchas a nerves or blood vessels.

FIG. 20A shows a diagram of apparatus 200 including a flexible probe214. The apparatus 200 includes a high voltage network 218, a lasersource 220, a probe assembly 214, and a target assembly 226. Accordingto one aspect of the invention, the apparatus 200 provides the requiredflexibility, without using strong magnetic fields, by locating electrongenerating and accelerating components in the target assembly 226. Theprobe assembly 214 couples both the laser source 220 and the highvoltage network 218 to the target assembly 226. The probe assemblyincludes flexible fiber optical cable 202 enclosed in a small-diameterflexible metallic tube 204.

The target assembly 226, which can be for example 1- to 2- cm in length,extends from the end of the probe assembly 214 and includes a shellwhich encloses the target 228. According to one embodiment, the targetassembly 226 is rigid in nature and generally cylindrical in shape. Inthis embodiment the cylindrical shell enclosing the target assembly canbe considered to provide a housing for the electron beam source as wellas a tubular probe extending from the housing along the electron beampath. The inner surface 226A of the assembly 226 is lined with anelectrical insulator, while the external surface 226b of the assembly226 is electrically conductive. According to a preferred embodiment, thetarget assembly is hermetically sealed to the end of the probe assembly214, and evacuated. According to another embodiment, the entire probeassembly 214 is evacuated.

The terminal end 202A of the fiber optical cable 202 is preferablycoated, over at least part of its area, with a semitransparentphotoemissive substance such as, Ag--O--Cs, thus forming a photocathode216. A high voltage conductor 208, embedded in the fiber optical cable202, conducts electrons to the cathode 216 from the high voltage network218. Similarly, the flexible tube 204 couples a ground return from thetarget 228 to the high voltage network 218, thereby establishing a highvoltage field between the cathode 216 and the target 228. The fiberoptical cable 202 acts as an insulating dielectric between the highvoltage conductor 208 and the grounded flexible tube 204.

In one embodiment, to eliminate absorption and scattering of the lightout of the fiber optic cable 202 by the high voltage wire 208, the fiberoptic cable 202 can have an annular configuration, as shown incross-section in FIG. 20B. The light from the laser 220 travels down theannular core 250 of the fiber optic cable 202. Cladding 260 on each sideof the core 250 has an index of refraction so as to refract the lightbeam incident on the interface back into the core 250. A groundedflexible metal tube 204 surrounds the outer cladding 260.

As in previously described embodiments, the target 228 can be forexample, beryllium; (Be), coated on one side with a thin film or layer228A of a higher atomic number element, such as tungsten (W) or gold(Au).

In operation, the small semiconductor laser 220 shining down the fiberoptical cable 202 activates the transmissive photocathode 216 whichgenerates free electrons 222. The high voltage field between the cathode216 and target 228 accelerates these electrons, thereby forcing them tostrike the surface 228A of target 228 and produce x-rays. In order togenerate, for example, 20 μA of current from an Ag--O--Cs photocathode216 with a laser 220 emitting light at a wavelength of 0.8 μm, the 0.4%quantum efficiency of this photocathode 216 for this wavelength requiresthat the laser 220 emits 7.5 mW optical power. Such diode lasers arereadily commercially available. According to the invention, thephotoemissive surface which forms cathode 216 can, in fact, be quitesmall. For example, for a current density at the cathode 216 of 1 A/cm²,the photoemitter's diameter need only be approximately 50 μm.

To minimize ion bombardment on the photocathode 216, a high electricalimpedance shield 217 is positioned in close proximity to thephotocathode 216 and electrically coupled along its outer edge, as shownin FIG. 22. The free electrons 222 can be directed through a smallaperture in the shield to disperse them over the target 228. Thereturning ions impinge on the shield 217 instead of on the photocathode216.

One difficult fabrication aspect of this invention is the fabrication ofthe photocathode 216, which for practical substances, with reasonablequantum efficiencies above 10⁻³, should be performed in a vacuum. Thisprocedure can be carried out with the fiber optical cable 202 positionedin a bell jar, where for example, an Ag--O--Cs photosurface isfabricated in the conventional manner. Subsequently, without exposure toair, the optical cable 202 can be inserted into the tube 204. The end202B can be vacuum sealed to the flexible tube 204.

In the above embodiments, the probe 14 or 214, along with its associatedtarget assembly 26, 126, or 226, can be coated with a biocompatibleouter layer, such as titanium nitride on a sublayer of nickel. Foradditional biocompatible protection a sheath of, for example,polyurethane can be fitted over the probe, such as that illustrated inFIG. 2.

The invention may be embodied in other specific forms without departingfrom the spirit or essential characteristics thereof. The presentembodiments are therefore to be considered in all respects asillustrative and not restrictive, the scope of the invention beingindicated by the appended claims rather than by the foregoingdescription, and all changes which come within the meaning and range ofequivalency of the claims are therefore intended to be embraced therein.

What is claimed is:
 1. An x-ray source comprising:A. a power supply,including a first terminal and a second terminal, and a drive means forestablishing an output voltage between said first terminal and saidsecond terminal, said output voltage having a peak value in theapproximate range of 10 kv to 90 kv; B. a flexible fiber optical cableassembly having an originating end and a terminating end, and includinga fiber optical element extending from said originating end to saidterminating end, and adapted for transmitting light incident on saidoriginating end to said terminating end; C. a light source, includingmeans for generating a beam of light at and directed to said originatingend of said fiber optical assembly; and D. a target assembly, affixed tosaid terminating end of said fiber optical cable assembly, andelectrically coupled to said power supply, by way of said first terminaland said second terminal and including means for emitting x-rays, in apredetermined spectral range, in response to light transmitted to saidterminating end.
 2. An x-ray source according to claim 1 wherein saidlight beam is substantially monochromatic.
 3. An x-ray source accordingto claim 2 where said light source is a laser and wherein said beam iscoherent.
 4. An x-ray source according to claim 1, wherein said targetassembly includes a photocathode having a photoemissive surface, saidphotocathode being positioned adjacent to said terminating end of saidfiber optical element and being responsive to portions of said lightbeam incident thereon from said terminating end to emit electrons fromsaid photoemissive surface.
 5. An x-ray source according to claim 4,wherein said target assembly includes a target element spaced apart fromand opposite said photoemissive surface, and including means foremitting x-rays in response to electrons incident on said target elementfrom said photoemissive surface.
 6. An x-ray source according to claim5, wherein said first terminal of said power supply is electricallycoupled to said photoemissive element and said second terminal of saidpower supply is electrically coupled to said target element, therebyestablishing an electric field which acts to accelerate electronsemitted from said photoemissive surface toward said target element. 7.An x-ray source according to claim 6, wherein said second terminal is atground potential.
 8. An x-ray source according to claim 6, wherein saidfiber optical cable assembly includes an electrical conductor locatedinternal to said fiber optical element, and adapted for electricallycoupling said first terminal of said power supply to said photocathode.9. An x-ray source according to claim 8, wherein said fiber opticalcable assembly includes an electrically conductive, flexible, outersheath, said sheath being adapted for electrically coupling said secondterminal of said power supply to said target assembly.
 10. An x-raysource according to claim 9, wherein said target assembly includes anelectrically conductive outer surface which couples between said sheathand said target element.
 11. An x-ray source according to claim 9,wherein said target assembly is substantially rigid in nature andgenerally cylindrical in shape, and includes an electrically insulatinginner surface, a first base end, and a second base end, wherein saidfirst base end opposes said second base end along a longitudinal axis,and wherein said photocathode is positioned proximate to said first baseend, and said target element is positioned proximate to said second baseend.
 12. An x-ray source according to claim 11, wherein said targetassembly includes means for sealing said target assembly to form aclosed chamber defined by said inner surface, said first base end, andsaid second base end.
 13. An x-ray source according to claim 12, whereinsaid closed chamber is evacuated.
 14. An x-ray source according to claim4, wherein said photocathode is formed on said terminating end of saidfiber optical element.
 15. An x-ray source according to claim 14 whereinsaid target assembly includes a target element, spaced apart from saidphotocathode, and including means for emitting x-rays in response toincident electrons.
 16. An x-ray source according to claim 15, whereinsaid first terminal of said power supply is electrically coupled to saidphotocathode and said second terminal of said power supply iselectrically coupled to said target element, thereby establishing anelectric field which acts to accelerate electrons emitted from saidphotoemissive surface toward said target element.
 17. An x-ray sourceaccording to claim 16, wherein said second terminal is at groundpotential.
 18. An x-ray source according to claim 16, wherein said fiberoptical cable assembly includes an electrical conductor located internalto said fiber optical element and adapted for coupling said firstterminal of said power supply to said photocathode.
 19. An x-ray sourceaccording to claim 18, wherein said fiber optical cable assemblyincludes an electrically conductive, flexible, outer sheath, said sheathbeing adapted for coupling said second terminal of said power supply tosaid target assembly.
 20. An x-ray source according to claim 19, whereinsaid target assembly includes and electrically conductive outer surfacewhich couples said sheath to said target element.
 21. An x-ray sourceaccording to claim 20, wherein said target assembly is substantiallyrigid in nature and generally cylindrical in shape, and includes anelectrically insulating inner surface, a first base end, and a secondbase end, wherein said first base end opposes said second base end alonga longitudinal axis, and wherein said photocathode is positionedproximate to said first base end, and said target element is positionedproximate to said second base end.
 22. An x-ray source according toclaim 1, wherein said power supply further includes selectively operablecontrol means for selectively controlling the amplitude of said outputvoltage.
 23. An x-ray source according to claim 5, wherein saidelectrons incident on said target element from said photoemissiveelement form a beam characterized by a current in the approximate rangeof 1 nA to 100 μA.
 24. An x-ray source according to claim 6, whereinsaid electrons incident on said target element from said photoemissivesurface are accelerated by said electric field to energies in theapproximate range of 10 keV to 90 keV.
 25. An x-ray source according toclaim 1, wherein said fiber optical cable assembly further comprises:A.an electrically conductive cable, wherein said fiber optical element isconcentrically disposed around said electrically conductive cable; andB. an electrically conductive outer shell, concentrically disposedaround said fiber optical element.
 26. An x-ray source according toclaim 25 wherein said fiber optical cable assembly further comprises afirst cladding shell, said first cladding shell having an index ofrefraction less than the index of refraction of said opticallytransmissive core and being concentrically disposed between saidelectrically conductive cable and said fiber optical element.
 27. Anx-ray source according to claim 26 wherein said fiber optical cableassembly further comprises a second cladding shell, said second claddingshell having an index of refraction less than the index of refraction ofsaid optically transmissive core and being concentrically disposedbetween said fiber optical element and said electrically conductiveouter shell.
 28. An x-ray source according to claim 5 further comprisinga toroidal shell shield element adjacent to said photocathode, saidshield element defining a central aperture permitting the passagetherethrough of certain of said emitted electrons to said target elementand blocking certain of the remainder of said emitted electrons.
 29. Anx-ray source according to claim 28 wherein said shield element is anelectrical high impedance material.